Bone treatments for injuries, genetic malformations and diseases often require implantation of grafts. It is well known that autografts and allografts are the safest implants; however, due to the limited supply and the risks of disease transmission and rejection encountered with these grafts, synthetic biomaterials have also been widely used as implants. Complications in vivo were observed with some of these biomaterials, as mechanical mismatches (stress shielding) and appearance of wear debris lead to bone atrophy, osteoporosis or osteolysis around the implants (Woo et al., 1976; Terjesen et al., 1988).
A new approach, defined as Tissue Engineering (TE), has recently raised a lot of interest. Tissue engineering involves the development of a new generation of biomaterials capable of specific interactions with biological tissues to yield functional tissue equivalents. The underlying concept is that cells can be isolated from a patient, expanded in cell culture and seeded onto a scaffold prepared from a specific biomaterial to form a scaffold/biological composite called a “TE construct”. The construct can then be grafted into the same patient to function as a replacement tissue. Some such systems are useful for organ tissue replacement where there is a limited availability of donor organs or where, in some cases (e.g. young patients) inadequate natural replacements are available. The scaffold itself may act as a delivery vehicle for biologically active moieties from growth factors, genes and drugs. This revolutionary approach to surgery has extensive applications with benefits to both patient well-being and the advancement of health care systems.
The application of tissue engineering to the growth of bone tissue involves harvesting osteogenic stem cells, seeding them and allowing them to grow to produce a new tissue in vitro. The newly obtained tissue can then be used as an autograft. Biodegradable polyesters—in particular poly(lactide-co-glycolide)s—have been used as scaffolds for tissue engineering of several different cell populations, for example: chondrocytes (as described by Freed et al. in the J. of Biomed. Mater. Res. 27:11-13,1993), hepatocytes (as described by Mooney et al. in the Journal of Biomedical Mat. Res. 29, 959-965, 1995) and most recently, bone marrow-derived cells (as described by Ishaug et al. in the J. Biomed. Mat. Res. 36: 17-28, 1997 and Holy et al., in Cells and Materials, 7, 223-234, 1997). Specifically, porous structures of these polyesters were prepared and seeded with cells; however, when bone marrow-derived cells were cultured on these porous structures, bone ingrowth only occurred within the outer edge of 3-D polymeric scaffold (Ishaug et al., supra; Holy et al., supra). Thus, the polymeric scaffolds prepared in these instances were inadequate to allow for the cell growth required to render tissue suitable for implantation or for use as an autograft.
The method of producing polymer scaffolds disclosed in Thomson et al., Fabrication of Biodegradable Polymer Scaffolds to Engineer Trabecular Bone”, J. Biomater. Sci. Polymer Edn. Vol. 7, No.1 pp. 23-38, 1995 VSP, involves formation of gelatin beads, after which a polymer is then “melted” at 80° C. and 333 g pressure around the beads after which the bead/polymer composite is cooled down, and the gelatin is leached out in distilled deionized water. The polymer is forming sheets of material around the beads and is in a solid state before the leaching of the beads/particulate.
U.S. Pat. No. 5,338,772 issued to Bauer et al. is directed to an implant material which is a composite of calcium phosphate ceramic particles and a bioadsorbable polymer. In the method of preparation disclosed in Bauer, calcium phosphate powder is mixed with a polymer and the mixture is subjected to microwave energy which melts the polymer to a liquid that forms a polymer coating around the particles with polymer bridges between encased particles.